Adjusting energy of a particle beam

ABSTRACT

An example particle accelerator includes a coil to provide a magnetic field to a cavity; a particle source to provide a plasma column to the cavity; a voltage source to provide a radio frequency (RF) voltage to the cavity to accelerate particles from the plasma column, where the magnetic field causes particles accelerated from the plasma column to move orbitally within the cavity; an enclosure containing an extraction channel to receive the particles accelerated from the plasma column and to output the received particles from the cavity; and a structure arranged proximate to the extraction channel to change an energy level of the received particles.

CROSS-REFERENCE TO RELATED APPLICATION

Priority is hereby claimed to U.S. Provisional Application No.61/707,515, which was filed on Sep. 28, 2012. The contents of U.S.Provisional Application No. 61/707,515 are hereby incorporated byreference into this disclosure.

TECHNICAL FIELD

This disclosure relates generally to adjusting energy of a particlebeam, such as a proton or ion beam used in a particle therapy system.

BACKGROUND

Particle therapy systems use an accelerator to generate a particle beamfor treating afflictions, such as tumors. In operation, the particlebeam is accelerated inside a cavity of the particle accelerator, andremoved from the cavity through an extraction channel. To track theextraction channel, the particle beam must be properly energized. Notenough energy, and the particle beam can collide with the inner edge ofthe extraction channel. Too much energy, and the particle beam cancollide with the outer edge of the extraction channel. The result isthat the particle beam is prevented from escaping the extractionchannel, or the portion of the particle beam that does escape iscompromised, thereby reducing treatment effectiveness.

Movement of the particle accelerator can affect the amount of energy inthe particle beam that is received at the extraction channel.

SUMMARY

An example particle accelerator includes a coil to provide a magneticfield to a cavity; a particle source to provide a plasma column to thecavity; a voltage source to provide a radio frequency (RF) voltage tothe cavity to accelerate particles from the plasma column, where themagnetic field causes particles accelerated from the plasma column tomove orbitally within the cavity; an enclosure containing an extractionchannel to receive the particles accelerated from the plasma column andto output the received particles from the cavity; and a structurearranged proximate to the extraction channel to change an energy levelof the received particles. This example particle accelerator may includeone or more of the following features, either alone or in combination.

The structure may have multiple thicknesses. The structure may havevariable thickness ranging from a maximum thickness to a minimumthickness. The structure may be movable relative to the extractionchannel to place one of the multiple thicknesses in a path of thereceived particles. The structure may be wheel-shaped and may berotatable within the extraction channel. The structure may include atleast one of the following materials: beryllium, carbon and plastic.

The particle accelerator may be rotatable relative to a fixed position.The particle accelerator may include a control system to controlmovement of the structure based on a rotational position of the particleaccelerator.

The particle accelerator may include a regenerator to adjust themagnetic field within the cavity to thereby change successive orbits ofthe particles accelerated from the plasma column so that, eventually,the particles output to the extraction channel.

An example proton therapy system may include the foregoing particleaccelerator, where the particles comprise protons; and a gantry on whichthe particle accelerator is mounted. The gantry is rotatable relative toa patient position. Protons are output essentially directly from theparticle accelerator to the patient position.

An example particle accelerator includes a coil to provide a magneticfield to a cavity; a particle source to provide a plasma column to thecavity; a voltage source to provide a radio frequency (RF) voltage tothe cavity to accelerate particles from the plasma column, where themagnetic field causes particles accelerated from the plasma column tomove orbitally within the cavity; an enclosure containing an extractionchannel to receive the particles accelerated from the plasma column andto output the received particles from the cavity; and a regenerator toadjust the magnetic field within the cavity to thereby change successiveorbits of the particles accelerated from the plasma column so that,eventually, the particles output to the extraction channel. Theregenerator is movable within the cavity relative to orbits of theparticles. This example particle accelerator may include one or more ofthe following features, either alone or in combination.

The regenerator may be configured to move radially relative to anapproximate center of the cavity. An actuator may be configured to movethe regenerator in response to a control signal. The particleaccelerator may be rotatable relative to a fixed position. The particleaccelerator may include a control system to generate the control signalto control movement of the regenerator based on a rotational position ofthe particle accelerator. The regenerator may include a ferromagneticmaterial, such as iron.

An example proton therapy system may include the foregoing particleaccelerator, where the particles comprise protons; and a gantry on whichthe particle accelerator is mounted. The gantry may be rotatablerelative to a patient position. Protons are output essentially directlyfrom the particle accelerator to the patient position.

An example particle accelerator includes a coil to provide a magneticfield to a cavity; a particle source to provide a plasma column to thecavity; a voltage source to provide a radio frequency (RF) voltage tothe cavity to accelerate particles from the plasma column, where themagnetic field causes particles accelerated from the plasma column tomove orbitally within the cavity; an enclosure containing an extractionchannel to receive the particles accelerated from the plasma column andto output the received particles from the cavity; and a regenerator toadjust the magnetic field within the cavity to thereby change successiveorbits of the particles accelerated from the plasma column so that,eventually, the particles output to the extraction channel. Theenclosure includes magnetic structures, where at least one of themagnetic structures has a slot therein, where the slot contains amagnetic shim that is ferromagnetic and movable within the slot, wherethe magnetic shim is movable relative to the regenerator to affect anamount by which the regenerator adjusts the magnetic field. This exampleparticle accelerator may include one or more of the following features,either alone or in combination.

The at least magnetic structure may have multiple slots therein. Eachslot may contain a magnetic shim that is ferromagnetic and that ismovable within the slot. Each magnetic shim may be movable relative tothe regenerator to affect an amount by which the regenerator adjusts themagnetic field.

The particle accelerator may be rotatable relative to a fixed position.The particle accelerator may include a control system to generate acontrol signal to control movement of the magnetic shim (or multiplemagnetic shims) based on a rotational position of the particleaccelerator. The magnetic shim (or multiple magnetic shims) may be orinclude an electromagnet.

An example proton therapy system may include the foregoing particleaccelerator, where the particles comprise protons; and a gantry on whichthe particle accelerator is mounted. The gantry is rotatable relative toa patient position. Protons are output essentially directly from theparticle accelerator to the patient position.

An example particle accelerator may include a cryostat comprising asuperconducting coil, where the superconducting coil conducts a currentthat generates a magnetic field; magnetic structures adjacent to thecryostat, where the cryostat is attached to the magnetic structures andthe magnetic structures contain a cavity; a particle source to provide aplasma column to the cavity; a voltage source to provide a radiofrequency (RF) voltage to the cavity to accelerate particles from theplasma column, where the magnetic field cause particles accelerated fromthe plasma column to move orbitally within the cavity; an extractionchannel to receive the particles accelerated from the plasma column andto output the received particles from the cavity; and an actuator thatis controllable to move the cryostat relative to the magneticstructures. This example particle accelerator may include one or more ofthe following features, either alone or in combination.

The particle accelerator may be rotatable relative to a fixed position.The particle accelerator may include a control system to generate acontrol signal to control the actuator based on a rotational position ofthe particle accelerator. The actuator may be controlled to controlmovement of the cryostat so as to compensate for effects of gravity onthe superconducting coil.

An example proton therapy system may include the foregoing particleaccelerator, where the particles comprise protons; and a gantry on whichthe particle accelerator is mounted. The gantry is rotatable relative toa patient position. Protons are output essentially directly from theparticle accelerator to the patient position.

An example variable-energy particle accelerator includes: magneticstructures defining a cavity in which particles are accelerated foroutput as a particle beam that has a selected energy from among a rangeof energies; an extraction channel to receive the particle beam; and astructure proximate to the extraction channel to intercept the particlebeam prior to the particle beam entering the extraction channel, wherethe structure is movable based on the selected energy, and where thestructure is for absorbing at least some energy of the particle beamprior to the particle beam entering the extraction channel. The examplevariable-energy particle accelerator may include one or more of thefollowing features, either alone or in combination.

The structure may be a wheel having varying thickness, where differentthicknesses are capable of absorbing different amounts of energy. Thevariable-energy particle accelerator may include a magnetic regeneratorto implement a magnetic field bump at a particle orbit that correspondsto the selected energy. The magnetic regenerator may be movable based onmovement of the variable-energy particle accelerator. The magneticregenerator may be movable to intercept a particle orbit having theselected energy.

Two or more of the features described in this disclosure, includingthose described in this summary section, may be combined to formimplementations not specifically described herein.

Control of the various systems described herein, or portions thereof,may be implemented via a computer program product that includesinstructions that are stored on one or more non-transitorymachine-readable storage media, and that are executable on one or moreprocessing devices. The systems described herein, or portions thereof,may be implemented as an apparatus, method, or electronic system thatmay include one or more processing devices and memory to storeexecutable instructions to implement control of the stated functions.

The details of one or more implementations are set forth in theaccompanying drawings and the description below. Other features,objects, and advantages will be apparent from the description anddrawings, and from the claims.

DESCRIPTION OF DRAWINGS

FIG. 1 is a perspective view of an example therapy system.

FIG. 2 is an exploded perspective view of components of an examplesynchrocyclotron.

FIGS. 3, 4, and 5 are cross-sectional views of an examplesynchrocyclotron.

FIG. 6 is a perspective view of an example synchrocyclotron.

FIG. 7 is a cross-sectional view of a portion of an example reversebobbin and windings.

FIG. 8 is a cross sectional view of an example cable-in-channelcomposite conductor.

FIG. 9 is a cross-sectional view of an example ion source.

FIG. 10 is a perspective view of an example dee plate and an exampledummy dee.

FIG. 11 is a perspective view of an example vault.

FIG. 12 is a perspective view of an example treatment room with a vault.

FIG. 13 shows an example of a patient relative to an accelerator.

FIG. 14 shows a patient positioned within an example inner gantry in atreatment room.

FIG. 15 is a top view of an example acceleration cavity and extractionchannel.

FIG. 16 is a graph showing magnetic field strength versus radialdistance from a plasma column, along with a cross-section of an examplepart of a cryostat of a superconducting magnet.

FIG. 17 is a top view of an example acceleration cavity and extractionchannel, which depicts orbits moving to enter the extraction channel.

FIG. 18 is a perspective view of an example structure used to change theenergy of a particle beam in the extraction channel.

FIG. 18A is a side view of the structure of FIG. 18.

FIGS. 19 to 21 are top views of an example acceleration cavity andextraction channel, which depict moving the regenerator to primarilyimpact certain orbits of particles in the cavity.

FIG. 22 is a perspective view of an example magnetic shim.

FIG. 23 is cut-away side view of magnetic yokes, an acceleration cavityand a cold mass, which includes magnetic shims.

FIG. 24 is a cut-away perspective view of an example part of a cryostat.

FIG. 25 is a conceptual view of an example particle therapy system thatmay use a variable-energy particle accelerator.

FIG. 26 is an example graph showing energy and current for variations inmagnetic field and distance in a particle accelerator.

FIG. 27 is a side view of an example structure for sweeping voltage on adee plate over a frequency range for each energy level of a particlebeam, and for varying the frequency range when the particle beam energyis varied.

FIG. 28 is a perspective, exploded view of an example magnet system thatmay be used in a variable-energy particle accelerator.

Like reference symbols in the various drawings indicate like elements.

DETAILED DESCRIPTION Overview

Described herein is an example of a particle accelerator for use in asystem, such as a proton or ion therapy system. The system includes aparticle accelerator—in this example, a synchrocyclotron—mounted on agantry. The gantry enables the accelerator to be rotated around apatient position, as explained in more detail below. In someimplementations, the gantry is steel and has two legs mounted forrotation on two respective bearings that lie on opposite sides of apatient. The particle accelerator is supported by a steel truss that islong enough to span a treatment area in which the patient lies and thatis attached stably at both ends to the rotating legs of the gantry. As aresult of rotation of the gantry around the patient, the particleaccelerator also rotates.

In an example implementation, the particle accelerator (e.g., thesynchrocyclotron) includes a cryostat that holds a superconducting coilfor conducting a current that generates a magnetic field (B). In thisexample, the cryostat uses liquid helium (He) to maintain the coil atsuperconducting temperatures, e.g., 4° Kelvin (K). Magnetic yokes areadjacent (e.g., around) the cryostat, and define a cavity in whichparticles are accelerated. The cryostat is attached to the magneticyokes through straps or the like. While this attachment, and theattachment of the superconducting coil inside the cryostat, restrictsmovement of the superconducting coil, coil movement is not entirelyprevented. For example, in some implementations, as a result ofgravitational pull during rotation of the gantry, the superconductingcoil is movable by small amounts (e.g., tenths of millimeters in somecases). As described below, this movement can affect the amount ofenergy in a particle beam that is received at an extraction channel andthereby affect the output of the particle accelerator.

In this example implementation, the particle accelerator includes aparticle source (e.g., a Penning Ion Gauge—PIG source) to provide aplasma column to the cavity. Hydrogen gas is ionized to produce theplasma column. A voltage source provides a radio frequency (RF) voltageto the cavity to accelerate particles from the plasma column. As noted,in this example, the particle accelerator is a synchrocyclotron.Accordingly, the RF voltage is swept across a range of frequencies toaccount for relativistic effects on the particles (e.g., increasingparticle mass) when extracting particles from the column. The magneticfield produced by the coil causes particles accelerated from the plasmacolumn to accelerate orbitally within the cavity. A magnetic fieldregenerator is positioned in the cavity to adjust the existing magneticfield inside the cavity to thereby change locations of successive orbitsof the particles accelerated from the plasma column so that, eventually,the particles output to an extraction channel that passes through theyokes. The regenerator may increase the magnetic field at a point in thecavity (e.g., it may produce a magnetic field “bump” at an area of thecavity), thereby causing each successive orbit of particles at thatpoint to precess outward toward the entry point of the extractionchannel, eventually reaching the extraction channel. The extractionchannel receives particles accelerated from the plasma column andoutputs the received particles from the cavity.

Movement of the superconducting coil can affect the locations of theorbits inside the cavity. For example, movement in one direction cancause lower-energy orbits to impact the regenerator, while movement inanother direction can cause higher-energy orbits to impact theregenerator (particle orbit energy is proportional to the radialdistance from the originating plasma column). So, in a case where overlylow-energy orbits impact the regenerator, the particle beam may collidewith the inner edge of the extraction channel, as noted above. In a casewhere overly high-energy orbits impact the regenerator, the particlebeam may collide with the outer edge of the extraction channel, as notedabove. The example systems described herein use techniques to compensatefor these effects resulting from motion of the superconducting coil dueto its rotation (e.g., due to the effect of gravity). A summary of thesetechniques is provided below, followed by a description of an exampleparticle therapy system in which they may be implemented and moredetailed descriptions of these various techniques.

In an example technique, a structure is incorporated proximate to (e.g.,at the entry to or inside of) the extraction channel. The structure maybe a rotatable variable-thickness wedge having a wheel-like shape. Thestructure absorbs energy of the particle beam, thereby allowing alower-energy (e.g., appropriately energized) beam to pass through theextraction channel. The thicker portions of the structure absorb moreenergy than the thinner portions of the structure. In someimplementations, the structure may contain no material at a point wherethe particle beam is meant to pass without any energy absorption.Alternatively, the structure may be movable out of the beam path. Thestructure thus enables the amount of energy in the beam to be variablyadjusted. In some implementations, the structure is controlled based ona rotational position of the particle accelerator. For example, theposition of the gantry may be determined, and that position may be usedto control the rotational position of the energy-absorbing structure.Ideally, the structure minimizes scattering of the beam; however, inpractice, there may be amounts of scatter that are present and that aretolerable.

In another example technique, the physical position of the regeneratorwithin the cavity may be adjustable to compensate for movement of thesuperconducting coil. For example, computer-controlled actuators may beused to adjust the position of the regenerator within the cavity based,e.g., on a rotational position of the particle accelerator. By soadjusting the position of the regenerator, it may be possible toposition the regenerator so that the appropriate adjustment to themagnetic field resulting from the regenerator impacts the properparticle orbits regardless of the rotational position of the particleaccelerator.

The regenerator is typically made of a ferromagnetic material. It istherefore possible to adjust the magnetic strength of the regeneratorusing one or more magnetic shims. Accordingly, in another exampletechnique, it is possible to adjust the magnetic field of theregenerator (e.g., to increase or decrease the magnetic field bumpproduced by the regenerator) or to move the effective location of themagnetic field perturbation produced by the regenerator withoutphysically moving the regenerator. For example, if movement of thesuperconducting coil results in lower-energy orbits impacting theregenerator, the magnetic field of the regenerator can be decreased sothat it doesn't begin to perturb beam orbits until higher energy orbitsreach it. It could also be effectively moved radially outward whilemaintaining the same overall strength (peak field) so that the orbitsgain higher energy before being effected by the regenerator. Likewise ifthe superconducting coil movement results in higher energy orbitsimpacting the regenerator the strength of the regenerator can beincreased or positioned radially inward to interact with orbits at lowerenergies. In an example implementation, the magnetic field is adjustedby moving a magnetic shim (e.g., a metal plunger) within a slot/hole ina magnetic yoke that is near to the regenerator. The magnetic shim ismade of ferromagnetic material and its proximity to the regeneratoraffects the magnetic field of the regenerator. Moving the magnetic shimcloser to the regenerator (e.g., further inside the slot increases themagnetic field produced by the regenerator; and moving the magnetic shimaway from the regenerator (e.g., upwards in, or outside of, the slot)decreases the magnetic field produced by the regenerator. In anotherexample the magnetic shim can be placed radially closer to the center ofthe cyclotron than the regenerator magnetic center. When the shim isplace closer to the acceleration plane, it moves the effective center ofthe regenerator magnetic perturbation without appreciably changing thepeak magnetic field strength. The magnetic shim may becomputer-controlled to vary its position based, e.g., on a rotationalposition of the particle accelerator.

In some implementations, more than one magnetic shim may be used. Instill other implementations, miniature electromagnet(s) may be used as amagnetic shim, and the current therethrough controlled based, e.g., on arotational position of the particle accelerator.

In another example, the entire cryostat may be moved relative to theyokes to compensate for movement of the superconducting coil. Forexample, movement of the cryostat can affect which orbits of particlesimpact the regenerator. So, if movement of the superconducting coiloccurs in one direction, the cryostat may be moved in the direction tocompensate for that movement and cause the superconducting coil to beproperly repositioned.

The foregoing techniques for adjusting the energy of a particle beam ina particle accelerator may be used individually in a single particleaccelerator, or any two or more of those techniques may be used in anyappropriate combination in a single particle accelerator. An example ofa particle therapy system in which the foregoing techniques may be usedis provided below.

Example Particle Therapy System

Referring to FIG. 1, a charged particle radiation therapy system 500includes a beam-producing particle accelerator 502 having a weight andsize small enough to permit it to be mounted on a rotating gantry 504with its output directed straight (that is, essentially directly) fromthe accelerator housing toward a patient 506.

In some implementations, the steel gantry has two legs 508, 510 mountedfor rotation on two respective bearings 512, 514 that lie on oppositesides of the patient. The accelerator is supported by a steel truss 516that is long enough to span a treatment area 518 in which the patientlies (e.g., twice as long as a tall person, to permit the person to berotated fully within the space with any desired target area of thepatient remaining in the line of the beam) and is attached stably atboth ends to the rotating legs of the gantry.

In some examples, the rotation of the gantry is limited to a range 520of less than 360 degrees, e.g., about 180 degrees, to permit a floor 522to extend from a wall of the vault 524 that houses the therapy systeminto the patient treatment area. The limited rotation range of thegantry also reduces the required thickness of some of the walls, whichprovide radiation shielding of people outside the treatment area. Arange of 180 degrees of gantry rotation is enough to cover all treatmentapproach angles, but providing a larger range of travel can be useful.For example the range of rotation may be between 180 and 330 degrees andstill provide clearance for the therapy floor space.

The horizontal rotational axis 532 of the gantry is located nominallyone meter above the floor where the patient and therapist interact withthe therapy system. This floor is positioned about 3 meters above thebottom floor of the therapy system shielded vault. The accelerator canswing under the raised floor for delivery of treatment beams from belowthe rotational axis. The patient couch moves and rotates in asubstantially horizontal plane parallel to the rotational axis of thegantry. The couch can rotate through a range 534 of about 270 degrees inthe horizontal plane with this configuration. This combination of gantryand patient rotational ranges and degrees of freedom allow the therapistto select virtually any approach angle for the beam. If needed, thepatient can be placed on the couch in the opposite orientation and thenall possible angles can be used.

In some implementations, the accelerator uses a synchrocyclotronconfiguration having a very high magnetic field superconductingelectromagnetic structure. Because the bend radius of a charged particleof a given kinetic energy is reduced in direct proportion to an increasein the magnetic field applied to it, the very high magnetic fieldsuperconducting magnetic structure permits the accelerator to be madesmaller and lighter. The synchrocyclotron uses a magnetic field that isuniform in rotation angle and falls off in strength with increasingradius. Such a field shape can be achieved regardless of the magnitudeof the magnetic field, so in theory there is no upper limit to themagnetic field strength (and therefore the resulting particle energy ata fixed radius) that can be used in a synchrocyclotron.

Superconducting materials lose their superconducting properties in thepresence of very high magnetic fields. High performance superconductingwire windings are used to allow very high magnetic fields to beachieved.

Superconducting materials typically need to be cooled to lowtemperatures for their superconducting properties to be realized. Insome examples described here, cryo-coolers are used to bring thesuperconducting coil windings to temperatures near absolute zero. Usingcryo-coolers can reduce complexity and cost.

The synchrocyclotron is supported on the gantry so that the beam isgenerated directly in line with the patient. The gantry permits rotationof the cyclotron about a horizontal rotational axis that contains apoint (isocenter 540) within, or near, the patient. The split truss thatis parallel to the rotational axis, supports the cyclotron on bothsides.

Because the rotational range of the gantry is limited, a patient supportarea can be accommodated in a wide area around the isocenter. Becausethe floor can be extended broadly around the isocenter, a patientsupport table can be positioned to move relative to and to rotate abouta vertical axis 542 through the isocenter so that, by a combination ofgantry rotation and table motion and rotation, any angle of beamdirection into any part of the patient can be achieved. The two gantryarms are separated by more than twice the height of a tall patient,allowing the couch with patient to rotate and translate in a horizontalplane above the raised floor.

Limiting the gantry rotation angle allows for a reduction in thethickness of at least one of the walls surrounding the treatment room.Thick walls, typically constructed of concrete, provide radiationprotection to individuals outside the treatment room. A wall downstreamof a stopping proton beam may be about twice as thick as a wall at theopposite end of the room to provide an equivalent level of protection.Limiting the range of gantry rotation enables the treatment room to besited below earth grade on three sides, while allowing an occupied areaadjacent to the thinnest wall reducing the cost of constructing thetreatment room.

In the example implementation shown in FIG. 1, the superconductingsynchrocyclotron 502 operates with a peak magnetic field in a pole gapof the synchrocyclotron of 8.8 Tesla. The synchrocyclotron produces abeam of protons having an energy of 250 MeV. In other implementationsthe field strength could be in the range of 4 to 20 Tesla or 6 to 20Tesla and the proton energy could be in the range of 150 to 300 MeV

The radiation therapy system described in this example is used forproton radiation therapy, but the same principles and details can beapplied in analogous systems for use in heavy ion (ion) treatmentsystems.

As shown in FIGS. 2, 3, 4, 5, and 6, an example synchrocyclotron 10(e.g., 502 in FIG. 1) includes a magnet system 12 that contains anparticle source 90, a radiofrequency drive system 91, and a beamextraction system 38. The magnetic field established by the magnetsystem has a shape appropriate to maintain focus of a contained protonbeam using a combination of a split pair of annular superconductingcoils 40, 42 and a pair of shaped ferromagnetic (e.g., low carbon steel)pole faces 44, 46.

The two superconducting magnet coils are centered on a common axis 47and are spaced apart along the axis. As shown in FIGS. 7 and 8, thecoils are formed by of Nb₃Sn-based superconducting 0.8 mm diameterstrands 48 (that initially comprise a niobium-tin core surrounded by acopper sheath) deployed in a twisted cable-in-channel conductorgeometry. After seven individual strands are cabled together, they areheated to cause a reaction that forms the final (brittle)superconducting material of the wire. After the material has beenreacted, the wires are soldered into the copper channel (outerdimensions 3.18×2.54 mm and inner dimensions 2.08×2.08 mm) and coveredwith insulation 52 (in this example, a woven fiberglass material). Thecopper channel containing the wires 53 is then wound in a coil having arectangular cross-section of 8.55 cm×19.02 cm, having 26 layers and 49turns per layer. The wound coil is then vacuum impregnated with an epoxycompound. The finished coils are mounted on an annular stainless steelreverse bobbin 56. Heater blankets 55 are placed at intervals in thelayers of the windings to protect the assembly in the event of a magnetquench.

The entire coil can then be covered with copper sheets to providethermal conductivity and mechanical stability and then contained in anadditional layer of epoxy. The precompression of the coil can beprovided by heating the stainless steel reverse bobbin and fitting thecoils within the reverse bobbin. The reverse bobbin inner diameter ischosen so that when the entire mass is cooled to 4 K, the reverse bobbinstays in contact with the coil and provides some compression. Heatingthe stainless steel reverse bobbin to approximately 50 degrees C. andfitting coils at a temperature of 100 degrees Kelvin can achieve this.

The geometry of the coil is maintained by mounting the coils in areverse rectangular bobbin 56 to exert a restorative force 60 that worksagainst the distorting force produced when the coils are energized. Asshown in FIG. 5, the coil position is maintained relative to the magnetyoke and cryostat using a set of warm-to-cold support straps 402, 404,406. Supporting the cold mass with thin straps reduces the heat leakageimparted to the cold mass by the rigid support system. The straps arearranged to withstand the varying gravitational force on the coil as themagnet rotates on board the gantry. They withstand the combined effectsof gravity and the large de-centering force realized by the coil when itis perturbed from a perfectly symmetric position relative to the magnetyoke. Additionally the links act to reduce dynamic forces imparted onthe coil as the gantry accelerates and decelerates when its position ischanged. Each warm-to-cold support includes one S2 fiberglass link andone carbon fiber link. The carbon fiber link is supported across pinsbetween the warm yoke and an intermediate temperature (50-70 K), and theS2 fiberglass link 408 is supported across the intermediate temperaturepin and a pin attached to the cold mass. Each link is 5 cm long (pincenter to pin center) and is 17 mm wide. The link thickness is 9 mm.Each pin is made of high strength stainless steel and is 40 mm indiameter.

Referring to FIG. 3, the field strength profile as a function of radiusis determined largely by choice of coil geometry and pole face shape;the pole faces 44, 46 of the permeable yoke material can be contoured tofine tune the shape of the magnetic field to ensure that the particlebeam remains focused during acceleration.

The superconducting coils are maintained at temperatures near absolutezero (e.g., about 4 degrees Kelvin) by enclosing the coil assembly (thecoils and the bobbin) inside an evacuated annular aluminum or stainlesssteel cryostatic chamber 70 that provides a free space around the coilstructure, except at a limited set of support points 71, 73. In analternate version (FIG. 4) the outer wall of the cryostat may be made oflow carbon steel to provide an additional return flux path for themagnetic field.

In some implementations, the temperature near absolute zero is achievedand maintained using one single-stage Gifford-McMahon cryo-cooler andthree two-stage Gifford McMahon cryo-coolers. Each two stage cryo-coolerhas a second stage cold end attached to a condenser that recondensesHelium vapor into liquid Helium. The cryo-cooler heads are supplied withcompressed Helium from a compressor. The single-stage Gifford-McMahoncryo-cooler is arranged to cool high temperature (e.g., 50-70 degreesKelvin) leads that supply current to the superconducting windings.

In some implementations, the temperature near absolute zero is achievedand maintained using two Gifford-McMahon cryo-coolers 72, 74 that arearranged at different positions on the coil assembly. Each cryo-coolerhas a cold end 76 in contact with the coil assembly. The cryo-coolerheads 78 are supplied with compressed Helium from a compressor 80. Twoother Gifford-McMahon cryo-coolers 77, 79 are arranged to cool hightemperature (e.g., 60-80 degrees Kelvin) leads that supply current tothe superconducting windings.

The coil assembly and cryostatic chambers are mounted within and fullyenclosed by two halves 81, 83 of a pillbox-shaped magnet yoke 82. Inthis example, the inner diameter of the coil assembly is about 74.6 cm.The iron yoke 82 provides a path for the return magnetic field flux 84and magnetically shields the volume 86 between the pole faces 44, 46 toprevent external magnetic influences from perturbing the shape of themagnetic field within that volume. The yoke also serves to decrease thestray magnetic field in the vicinity of the accelerator. In someimplementations, the synchrocyclotron may have an active return systemto reduce stray magnetic fields. An example of an active return systemis described in U.S. patent application Ser. No. 13/907,601, which wasfiled on May 31, 2013, the contents of which are incorporated herein byreference. In the active return system, the relatively large magneticyokes described herein are replaced by smaller magnetic structures,referred to as pole pieces. Superconducting coils run current oppositeto the main coils described herein in order to provide magnetic returnand thereby reduce stray magnetic fields

As shown in FIGS. 3 and 9, the synchrocyclotron includes a particlesource 90 of a Penning ion gauge geometry located near the geometriccenter 92 of the magnet structure 82. The particle source may be asdescribed below, or the particle source may be of the type described inU.S. patent application Ser. No. 11/948,662 incorporated herein byreference.

Particle source 90 is fed from a supply 99 of hydrogen through a gasline 101 and tube 194 that delivers gaseous hydrogen. Electric cables 94carry an electric current from a current source 95 to stimulate electrondischarge from cathodes 192, 190 that are aligned with the magneticfield, 200.

In some implementations, the gas in gas tube 101 may include a mixtureof hydrogen and one or more other gases. For example, the mixture maycontain hydrogen and one or more of the noble gases, e.g., helium, neon,argon, krypton, xenon and/or radon (although the mixture is not limitedto use with the noble gases). In some implementations, the mixture maybe a mixture of hydrogen and helium. For example, the mixture maycontain about 75% or more of hydrogen and about 25% or less of helium(with possible trace gases included). In another example, the mixturemay contain about 90% or more of hydrogen and about 10% or less ofhelium (with possible trace gases included). In examples, thehydrogen/helium mixture may be any of thefollowing: >95%/<5%, >90%/<10%, >85%/<15%, >80%/<20%, >75%/<20%, and soforth.

Possible advantages of using a noble (or other) gas in combination withhydrogen in the particle source may include: increased beam intensity,increased cathode longevity, and increased consistency of beam output.

In this example, the discharged electrons ionize the gas exiting througha small hole from tube 194 to create a supply of positive ions (protons)for acceleration by one semicircular (dee-shaped) radio-frequency plate100 that spans half of the space enclosed by the magnet structure andone dummy dee plate 102. In the case of an interrupted particle source(an example of which is described in U.S. patent application Ser. No.11/948,662), all (or a substantial part) of the tube containing plasmais removed at the acceleration region, thereby allowing ions to be morerapidly accelerated in a relatively high magnetic field.

As shown in FIG. 10, the dee plate 100 is a hollow metal structure thathas two semicircular surfaces 103, 105 that enclose a space 107 in whichthe protons are accelerated during half of their rotation around thespace enclosed by the magnet structure. A duct 109 opening into thespace 107 extends through the yoke to an external location from which avacuum pump 111 can be attached to evacuate the space 107 and the restof the space within a vacuum chamber 119 in which the acceleration takesplace. The dummy dee 102 comprises a rectangular metal ring that isspaced near to the exposed rim of the dee plate. The dummy dee isgrounded to the vacuum chamber and magnet yoke. The dee plate 100 isdriven by a radio-frequency signal that is applied at the end of aradio-frequency transmission line to impart an electric field in thespace 107. The radio frequency electric field is made to vary in time asthe accelerated particle beam increases in distance from the geometriccenter. The radio frequency electric field may be controlled in themanner described in U.S. patent application Ser. No. 11/948,359,entitled “Matching A Resonant Frequency Of A Resonant Cavity To AFrequency Of An Input Voltage”, the contents of which are incorporatedherein by reference.

For the beam emerging from the centrally located particle source toclear the particle source structure as it begins to spiral outward, alarge voltage difference is required across the radio frequency plates.20,000 Volts is applied across the radio frequency plates. In someversions from 8,000 to 20,000 Volts may be applied across the radiofrequency plates. To reduce the power required to drive this largevoltage, the magnet structure is arranged to reduce the capacitancebetween the radio frequency plates and ground. This is done by formingholes with sufficient clearance from the radio frequency structuresthrough the outer yoke and the cryostat housing and making sufficientspace between the magnet pole faces.

The high voltage alternating potential that drives the dee plate has afrequency that is swept downward during the accelerating cycle toaccount for the increasing relativistic mass of the protons and thedecreasing magnetic field. The dummy dee does not require a hollowsemi-cylindrical structure as it is at ground potential along with thevacuum chamber walls. Other plate arrangements could be used such asmore than one pair of accelerating electrodes driven with differentelectrical phases or multiples of the fundamental frequency. The RFstructure can be tuned to keep the Q high during the required frequencysweep by using, for example, a rotating capacitor having intermeshingrotating and stationary blades. During each meshing of the blades, thecapacitance increases, thus lowering the resonant frequency of the RFstructure. The blades can be shaped to create a precise frequency sweeprequired. A drive motor for the rotating condenser can be phase lockedto the RF generator for precise control. One bunch of particles isaccelerated during each meshing of the blades of the rotating condenser.

The vacuum chamber 119 in which the acceleration occurs is a generallycylindrical container that is thinner in the center and thicker at therim. The vacuum chamber encloses the RF plates and the particle sourceand is evacuated by the vacuum pump 111. Maintaining a high vacuuminsures that accelerating ions are not lost to collisions with gasmolecules and enables the RF voltage to be kept at a higher levelwithout arcing to ground.

Protons traverse a generally spiral orbital path beginning at theparticle source. In half of each loop of the spiral path, the protonsgain energy as they pass through the RF electric field in space 107. Asthe ions gain energy, the radius of the central orbit of each successiveloop of their spiral path is larger than the prior loop until the loopradius reaches the maximum radius of the pole face. At that location amagnetic and electric field perturbation directs ions into an area wherethe magnetic field rapidly decreases, and the ions depart the area ofthe high magnetic field and are directed through an evacuated tube 38,referred to herein as the extraction channel, to exit the yoke of thecyclotron. A magnetic regenerator may be used to change the magneticfield perturbation to direct the ions. The ions exiting the cyclotronwill tend to disperse as they enter the area of markedly decreasedmagnetic field that exists in the room around the cyclotron. Beamshaping elements 107, 109 in the extraction channel 38 redirect the ionsso that they stay in a straight beam of limited spatial extent.

The magnetic field within the pole gap needs to have certain propertiesto maintain the beam within the evacuated chamber as it accelerates. Themagnetic field index n, which is shown below,

n=−(r/B)dB/dr,

should be kept positive to maintain this “weak” focusing. Here r is theradius of the beam and B is the magnetic field. Additionally, in someimplementations, the field index needs to be maintained below 0.2,because at this value the periodicity of radial oscillations andvertical oscillations of the beam coincide in a vr=2 v_(z) resonance.The betatron frequencies are defined by v_(r)=(1−n)^(1/2) andv_(z)=n^(1/2). The ferromagnetic pole face is designed to shape themagnetic field generated by the coils so that the field index n ismaintained positive and less than 0.2 in the smallest diameterconsistent with a 250 MeV beam in the given magnetic field.

As the beam exits the extraction channel it is passed through a beamformation system 125 (FIG. 5) that can be programmably controlled tocreate a desired combination of scattering angle and range modulationfor the beam. Beam formation system 125 may be used in conjunction withan inner gantry 601 (FIG. 14) to direct a beam to the patient.

During operation, the plates absorb energy from the applied radiofrequency field as a result of conductive resistance along the surfacesof the plates. This energy appears as heat and is removed from theplates using water cooling lines 108 that release the heat in a heatexchanger 113 (FIG. 3).

Stray magnetic fields exiting from the cyclotron are limited by both thepillbox magnet yoke (which also serves as a shield) and a separatemagnetic shield 114. The separate magnetic shield includes of a layer117 of ferromagnetic material (e.g., steel or iron) that encloses thepillbox yoke, separated by a space 116. This configuration that includesa sandwich of a yoke, a space, and a shield achieves adequate shieldingfor a given leakage magnetic field at lower weight.

As mentioned, the gantry allows the synchrocyclotron to be rotated aboutthe horizontal rotational axis 532. The truss structure 516 has twogenerally parallel spans 580, 582. The synchrocyclotron is cradledbetween the spans about midway between the legs. The gantry is balancedfor rotation about the bearings using counterweights 122, 124 mounted onends of the legs opposite the truss.

The gantry is driven to rotate by an electric motor mounted to one orboth of the gantry legs and connected to the bearing housings by drivegears. The rotational position of the gantry is derived from signalsprovided by shaft angle encoders incorporated into the gantry drivemotors and the drive gears.

At the location at which the ion beam exits the cyclotron, the beamformation system 125 acts on the ion beam to give it properties suitablefor patient treatment. For example, the beam may be spread and its depthof penetration varied to provide uniform radiation across a given targetvolume. The beam formation system can include passive scatteringelements as well as active scanning elements.

All of the active systems of the synchrocyclotron (the current drivensuperconducting coils, the RF-driven plates, the vacuum pumps for thevacuum acceleration chamber and for the superconducting coil coolingchamber, the current driven particle source, the hydrogen gas source,and the RF plate coolers, for example), may be controlled by appropriatesynchrocyclotron control electronics (not shown), which may include,e.g., one or more computers programmed with appropriate programs toeffect control.

The control of the gantry, the patient support, the active beam shapingelements, and the synchrocyclotron to perform a therapy session isachieved by appropriate therapy control electronics (not shown).

As shown in FIGS. 1, 11, and 12, the gantry bearings are supported bythe walls of a cyclotron vault 524. The gantry enables the cyclotron tobe swung through a range 520 of 180 degrees (or more) includingpositions above, to the side of, and below the patient. The vault istall enough to clear the gantry at the top and bottom extremes of itsmotion. A maze 146 sided by walls 148, 150 provides an entry and exitroute for therapists and patients. Because at least one wall 152 is notin line with the proton beam directly from the cyclotron, it can be maderelatively thin and still perform its shielding function. The otherthree side walls 154, 156, 150/148 of the room, which may need to bemore heavily shielded, can be buried within an earthen hill (not shown).The required thickness of walls 154, 156, and 158 can be reduced,because the earth can itself provide some of the needed shielding.

Referring to FIGS. 12 and 13, for safety and aesthetic reasons, atherapy room 160 may be constructed within the vault. The therapy roomis cantilevered from walls 154, 156, 150 and the base 162 of thecontaining room into the space between the gantry legs in a manner thatclears the swinging gantry and also maximizes the extent of the floorspace 164 of the therapy room. Periodic servicing of the accelerator canbe accomplished in the space below the raised floor. When theaccelerator is rotated to the down position on the gantry, full accessto the accelerator is possible in a space separate from the treatmentarea. Power supplies, cooling equipment, vacuum pumps and other supportequipment can be located under the raised floor in this separate space.Within the treatment room, the patient support 170 can be mounted in avariety of ways that permit the support to be raised and lowered and thepatient to be rotated and moved to a variety of positions andorientations.

In system 602 of FIG. 14, a beam-producing particle accelerator of thetype described herein, in this case synchrocyclotron 604, is mounted onrotating gantry 605. Rotating gantry 605 is of the type describedherein, and can angularly rotate around patient support 606. Thisfeature enables synchrocyclotron 604 to provide a particle beam directlyto the patient from various angles. For example, as in FIG. 14, ifsynchrocyclotron 604 is above patient support 606, the particle beam maybe directed downwards toward the patient. Alternatively, ifsynchrocyclotron 604 is below patient support 606, the particle beam maybe directed upwards toward the patient. The particle beam is applieddirectly to the patient in the sense that an intermediary beam routingmechanism is not required. A routing mechanism, in this context, isdifferent from a shaping or sizing mechanism in that a shaping or sizingmechanism does not re-route the beam, but rather sizes and/or shapes thebeam while maintaining the same general trajectory of the beam.

Further details regarding an example implementation of the foregoingsystem may be found in U.S. Pat. No. 7,728,311, filed on Nov. 16, 2006and entitled “Charged Particle Radiation Therapy”, and in U.S. patentapplication Ser. No. 12/275,103, filed on Nov. 20, 2008 and entitled“Inner Gantry”. The contents of U.S. Pat. No. 7,728,311 and in U.S.patent application Ser. No. 12/275,103 are hereby incorporated byreference into this disclosure. In some implementations, thesynchrocyclotron may be a variable-energy device, such as that describedin U.S. patent application Ser. No. 13/916,401, filed on Jun. 12, 2013,the contents of which are incorporated herein by reference.

Example Implementations

FIG. 15 shows a top view of a portion of a cavity 700 in which particlesare accelerated orbitally (e.g., in outward spiral orbits). A particlesource 701, examples of which are described above, is disposed at aboutthe center of the cavity. Charged particles (e.g., protons or ions) areextracted from a plasma column generated by particle source 701. Thecharged particles accelerate outwardly in orbits toward, and eventuallyreaching, magnetic regenerator 702. In this example implementation,regenerator 702 is a ferromagnetic structure made, e.g., of steel, iron,or any other type of ferromagnetic material. Regenerator 702 alters thebackground magnetic field that causes the outward orbital acceleration.In this example, regenerator 702 augments that magnetic field (e.g., itprovides a bump in the field). The bump in the background magnetic fieldaffects the particle orbits in a way that causes the orbits to moveoutwardly towards extraction channel 703. Eventually, the orbits enterextraction channel 703, from which they exit.

In more detail, a particle beam orbit approaches and interacts withregenerator 702. As a result of the increased magnetic field, theparticle beam turns a bit more there and, instead of being circular, itprecesses to the extraction channel. FIG. 16 shows the magnetic field(B) plotted against the radius (r) relative to the particle source 702.As shown in FIG. 16, in this example, B varies from about 9 Tesla (T) toabout −2T. The 9T occurs at about the center of cavity 700. The polarityof the magnetic field changes after the magnetic field crosses thesuperconducting coil, resulting in about −2T on the exterior of thecoil, eventually fading to about zero. The magnetic field bump 705occurs at the point of the regenerator. FIG. 16 also shows the magneticfield plot relative to a cross-section 706 of a bobbin 706 havingextraction channel 703 between two superconducting coils 709, 710.

Referring to FIG. 17, regenerator 702 causes changes in the angle andpitch of orbits 710 so that they move toward extraction channel 703. Atthe point of the extraction channel, the magnetic field strength issufficiently low to enable the particle beam to enter the extractionchannel and to proceed therethrough. Referring back to FIG. 15,extraction channel 703 contains various magnetic structures 711 foradding and/or subtracting dipole fields to direct the entering particlebeam through extraction channel 703, to beam shaping elements.

In order to reach the exit point, the particle beam should have theappropriate amount of energy. The amount of energy required to reachthat point may vary based, e.g., on the size of the accelerator and thelength of the extraction channel (in this example, the extractionchannel is about 1.7 or 2 meters in length). In this regard, at leastpart of extraction channel 703 is above the superconducting coil. Assuch, the magnetic field in the extraction channel may change little inresponse to accelerator rotation. Accordingly, the amount of energyneeded for a particle beam to traverse the extraction channel may notchange appreciably in response to the rotation of the particleaccelerator.

As explained above, as the superconducting coil moves during rotation,orbits that are affected by regenerator 702 change due to gravitationalmovement of the coil. As noted, this movement can be as little as tenthsof millimeters. Nevertheless, as a result, the energy of the particlebeam that enters the extraction channel may be different from the energyrequired to traverse the entire channel. To adjust for this change inthe energy of particles entering the extraction channel, a structure 715may be placed inside, or at the entry point to, extraction channel 703.The structure may be used to absorb excess energy in the particle beam.In this example, structure 715 is a rotatable, variable-thickness wedge,which may have a wheel-like shape. An example of structure 715 is shownin FIGS. 18 and 18A. As shown in these figures, structure 715 may havecontinuously varying thickness. Alternatively, the thicknesses may varystep-wise.

The structure may be moved (e.g., rotated) to absorb an appropriateamount of energy from a particle beam in/entering the extractionchannel. In this implementation, thicker parts 715 a of the structureabsorb more energy than thinner parts 715 b. Accordingly, the structuremay be moved (e.g., rotated) to absorb different amounts of energy in aparticle beam. In some implementations, the structure may have a partcontaining no material (e.g., a “zero” thickness), which allows theparticle beam to pass unaltered. Alternatively, in such cases, thestructure may be moved entirely or partly out of the beam path. In someimplementations, the maximum thickness may be on the order ofcentimeters; however, the maximum thickness will vary fromsystem-to-system based, e.g., on energy absorbing requirements. FIG. 18Aalso shows a motor 716 that controls an axle to rotate structure 715,e.g., in response to a detected gantry position.

The structure may be made of any appropriate material that is capable ofabsorbing energy in a particle beam. As noted above, ideally, thestructure minimizes scattering of the particle beam in the extractionchannel; however, in practice, there may be amounts of scatter that arepresent and that are tolerable. Examples of materials that may be usedfor the structure include, but are not limited to, beryllium, plasticcontaining hydrogen, and carbon. These materials may be used alone, incombination, or in combination with other materials.

The movement (e.g., rotation) of the structure may becomputer-controlled using a control system that is part of the broaderparticle therapy system. Computer control may include generating one ormore control signals to control movement of mechanical devices, such asactuators and motors that produce the motion. The rotation of structure715 may be controlled based on a rotational position of the particleaccelerator, as measured by the rotational position of the gantry (see,e.g., FIGS. 1, 11 and 12 showing gantry rotation) on which the particleaccelerator is mounted. The various parameters used to set therotational position of the structure vis-à-vis the position of thegantry may be measured empirically, and programmed into the controlsystem computer.

As noted above, in some implementations, the magnetic field in theextraction channel may change, albeit very little, in response toaccelerator rotation. The amount of the change may be, e.g., a fewtenths of a percent. In a specific example, this is reflected by achange of about six amperes (amps) of current out of a normal ˜2000 ampsrunning through the superconducting coil. This can affect the energyrequired for a particle beam to traverse the extraction channel. Thissmall change in magnetic field may be adjusted by controlling thecurrent through the superconducting coil or by controlling the rotationof structure 715.

In other implementations, adjusting the energy of particle beamsreaching the extraction channel may be achieved by physically movingregenerator 702 so that, at different rotational positions, theregenerator affects different particle orbits. As above, the movement orregenerator 702 may be computer-controlled through a control system thatis part of the particle therapy system. For example, the movement ofregenerator 702 may be controlled based on a rotational position of theparticle accelerator, as measured by the rotational position of thegantry on which the particle accelerator is mounted. The variousparameters used to set the location of the regenerator vis-à-vis therotational position of the gantry may be measured empirically, andprogrammed into the control system computer. One or morecomputer-controlled actuators may effect actual movement of theregenerator.

Referring to FIG. 19 for example, regenerator 702 may be positioned atlocation 717 initially, e.g., at a predefined initial position of theaccelerator. In this position, the magnetic field bump produced by theregenerator has a primary impact on orbit 719 (to direct particles atthat orbital position to the extraction channel). Orbit 720 is furtherfrom the location 721 of the plasma column than orbit 719. Consequently,orbit 720 has higher energy than orbit 719. Orbit 722 is closer to thelocation 721 of the plasma column than orbit 719. Consequently, orbit722 has lower energy than orbit 719. As shown in FIG. 20, movement ofthe superconducting coil as a result of rotation can cause lower-energyorbit 722 to move into the path of regenerator 702 such that regenerator702 primarily affects orbit 722. However, because orbit 722 is lowerenergy, it may not be capable of traversing the extraction channel andmay impact the inner wall of the extraction channel before exiting.Accordingly, regenerator 702 may be moved from location 717 to location723 (as depicted by arrow 724 of FIG. 21) so that regenerator 702 againprimarily impacts orbit 719. The converse may be true as well. That is,if the superconducting coil moves such that an overly high-energy orbit720 is primarily impacted by regenerator 702, regenerator 702 may bemoved in the other direction (e.g., towards location 721) so that itprimarily impacts the lower-energy orbit 719 (which has also moved).Although the figures depict movement of the regenerator in one dimension(radially), the regenerator may be moved in two or three dimensions,e.g., it may be moved in the Cartesian X, Y and/or Z directions.

In other implementations, the orbit that is affected primarily by theregenerator may be changed by altering the magnetic field (the magneticfield bump). This may be done, e.g., by changing the amount offerromagnetic material in proximity to the regenerator. In animplementation, one or more magnetic shims may be used to alter theshape and/or strength of the magnetic field produced by the regenerator.In this regard, the regenerator may be made of a ferromagnetic material,such as steel (although other materials may be used in place of, or inaddition to, steel). The magnetic shims may be a ferromagnetic materialthat is different from, or the same as, the material of which theregenerator is made.

In this implementation, the magnetic shims includes one or iron or steelmagnetic shims. An example is magnetic shim 730 shown in FIG. 22;however, any appropriate shape may be used. For example, magnetic shim730 may be in the shape of a rod or may have other appropriate shapes.Referring to FIG. 23, magnetic shims 730 a, 730 b may be placed in aslot of the corresponding yoke 731 a, 731 b near to the regenerator 702or in the regenerator itself. Moving the magnetic shim downward, furtherinside a slot in the yoke increases the amount of ferromagnetic materialnear to the regenerator and thereby alters the location and size of themagnetic field bump produced by the regenerator. By contrast, moving amagnetic shim upward and out of the yoke decreases the amount offerromagnetic material near to the regenerator and thereby alters thelocation and size of the magnetic field bump produced by theregenerator. Increasing the amount of ferromagnetic material causes themagnetic field bump to be moved inward (towards the plasma column—see,e.g., FIGS. 19 to 21) to thereby primarily affect lower-energy particleorbits. Decreasing the amount of ferromagnetic material causes themagnetic field bump to be moved outward (away the plasma column) tothereby primarily affect higher-energy particle orbits.

The magnetic shims may be permanently screwed into the yoke and held inplace there using screws or they may be controlled in real time. In thisregard, movement or the magnetic shim(s) may be computer-controlledthrough a control system that is part of the particle therapy system.For example, the movement of each magnetic shim 730 a, 730 b may becontrolled based on a rotational position of the particle accelerator,as measured by the rotation position of the gantry on which the particleaccelerator is mounted. The various parameters used to set the magneticshim location vis-à-vis the rotational position of the accelerator maybe measured empirically, and programmed into the control systemcomputer. One or more computer-controlled actuators may effect actualmovement of the magnetic shim(s). Although only two magnetic shims aredepicted, any number of magnetic shims may be used (e.g., one or more).

In some implementations, the magnetic shim(s) (e.g., the magneticshim(s) described above) may instead be, or include, one or moreminiature electromagnets, the current through which is controlled toaffect the magnetic field produced by the regenerator in the mannerdescribed above. The current through the one or more electromagnets maybe computer-controlled through a control system that is part of theparticle therapy system. For example, the current may be controlledbased on a rotational position of the particle accelerator, as measuredby the rotation position of the gantry on which the particle acceleratoris mounted. The various parameters used to set the current vis-à-vis therotational position of the accelerator may be measured empirically, andprogrammed into the control system computer.

In other implementations, adjusting the energy of particle beamsreaching the extraction channel may be achieved by physically moving thecryostat to compensate for movement of the coil as a result of rotation.For example, the cryostat may be moved in a direction opposite to thedirection that the coil moves. As above, the movement of the cryostatmay be computer-controlled through a control system that is part of theparticle therapy system. For example, the movement of cryostat may becontrolled based on a rotational position of the particle accelerator,as measured by the rotation position of the gantry on which the particleaccelerator is mounted. The various parameters used to set the movementof the cryostat vis-à-vis the rotational position of the gantry may bemeasured empirically, and programmed into the control system computer.One or more computer-controlled actuators may effect actual movement ofthe cryostat.

Referring to FIG. 24, for example, rotation of the accelerator may causecoils 709, 710 to move in the direction of arrow 735 within theirrespective chambers. In response, the position of cryostat 736 may bechanged, e.g., cryostat 736 may be moved, e.g., in the direction ofarrow 737 (e.g., in the opposite direction by an opposite amount). Thismovement causes a corresponding movement of coil 709, 710, therebybringing coils 709, 710 back into their original position in properalignment relative to the regenerator.

Variable-Energy Particle Accelerator

The particle accelerator used in the example particle therapy systemsdescribed herein may be a variable-energy particle accelerator.

The energy of the extracted particle beam (the particle beam output fromthe accelerator) can affect the use of the particle beam duringtreatment. In some machines, the energy of the particle beam (orparticles in the particle beam) does not increase after extraction.However, the energy may be reduced based on treatment needs after theextraction and before the treatment. Referring to FIG. 25, an exampletreatment system 910 includes an accelerator 912, e.g., asynchrocyclotron, from which a particle (e.g., proton) beam 914 having avariable energy is extracted to irradiate a target volume 924 of a body922. Optionally, one or more additional devices, such as a scanning unit916 or a scattering unit 916, one or more monitoring units 918, and anenergy degrader 920, are placed along the irradiation direction 928. Thedevices intercept the cross-section of the extracted beam 914 and alterone or more properties of the extracted beam for the treatment.

A target volume to be irradiated (an irradiation target) by a particlebeam for treatment typically has a three-dimensional configuration. Insome examples, to carry-out the treatment, the target volume is dividedinto layers along the irradiation direction of the particle beam so thatthe irradiation can be done on a layer-by-layer basis. For certain typesof particles, such as protons, the penetration depth (or which layer thebeam reaches) within the target volume is largely determined by theenergy of the particle beam. A particle beam of a given energy does notreach substantially beyond a corresponding penetration depth for thatenergy. To move the beam irradiation from one layer to another layer ofthe target volume, the energy of the particle beam is changed.

In the example shown in FIG. 25, the target volume 924 is divided intonine layers 926 a-926 i along the irradiation direction 928. In anexample process, the irradiation starts from the deepest layer 926 i,one layer at a time, gradually to the shallower layers and finishes withthe shallowest layer 926 a. Before application to the body 922, theenergy of the particle beam 914 is controlled to be at a level to allowthe particle beam to stop at a desired layer, e.g., the layer 926 d,without substantially penetrating further into the body or the targetvolume, e.g., the layers 926 e-926 i or deeper into the body. In someexamples, the desired energy of the particle beam 914 decreases as thetreatment layer becomes shallower relative to the particle acceleration.In some examples, the beam energy difference for treating adjacentlayers of the target volume 924 is about 3 MeV to about 100 MeV, e.g.,about 10 MeV to about 80 MeV, although other differences may also bepossible, depending on, e.g., the thickness of the layers and theproperties of the beam.

The energy variation for treating different layers of the target volume924 can be performed at the accelerator 912 (e.g., the accelerator canvary the energy) so that, in some implementations, no additional energyvariation is required after the particle beam is extracted from theaccelerator 912. So, the optional energy degrader 920 in the treatmentsystem 10 may be eliminated from the system. In some implementations,the accelerator 912 can output particle beams having an energy thatvaries between about 100 MeV and about 300 MeV, e.g., between about 115MeV and about 250 MeV. The variation can be continuous ornon-continuous, e.g., one step at a time. In some implementations, thevariation, continuous or non-continuous, can take place at a relativelyhigh rate, e.g., up to about 50 MeV per second or up to about 20 MeV persecond. Non-continuous variation can take place one step at a time witha step size of about 10 MeV to about 90 MeV.

When irradiation is complete in one layer, the accelerator 912 can varythe energy of the particle beam for irradiating a next layer, e.g.,within several seconds or within less than one second. In someimplementations, the treatment of the target volume 924 can be continuedwithout substantial interruption or even without any interruption. Insome situations, the step size of the non-continuous energy variation isselected to correspond to the energy difference needed for irradiatingtwo adjacent layers of the target volume 924. For example, the step sizecan be the same as, or a fraction of, the energy difference.

In some implementations, the accelerator 912 and the degrader 920collectively vary the energy of the beam 914. For example, theaccelerator 912 provides a coarse adjustment and the degrader 920provides a fine adjustment or vice versa. In this example, theaccelerator 912 can output the particle beam that varies energy with avariation step of about 10-80 MeV, and the degrader 920 adjusts (e.g.,reduces) the energy of the beam at a variation step of about 2-10 MeV.

The reduced use (or absence) of the energy degrader, which can includerange shifters, helps to maintain properties and quality of the outputbeam from the accelerator, e.g., beam intensity. The control of theparticle beam can be performed at the accelerator. Side effects, e.g.,from neutrons generated when the particle beam passes the degrader 920can be reduced or eliminated.

The energy of the particle beam 914 may be adjusted to treat anothertarget volume 930 in another body or body part 922′ after completingtreatment in target volume 924. The target volumes 924, 930 may be inthe same body (or patient), or may belong to different patients. It ispossible that the depth D of the target volume 930 from a surface ofbody 922′ is different from that of the target volume 924. Although someenergy adjustment may be performed by the degrader 920, the degrader 912may only reduce the beam energy and not increase the beam energy.

In this regard, in some cases, the beam energy required for treatingtarget volume 930 is greater than the beam energy required to treattarget volume 924. In such cases, the accelerator 912 may increase theoutput beam energy after treating the target volume 924 and beforetreating the target volume 930. In other cases, the beam energy requiredfor treating target volume 930 is less than the beam energy required totreat target volume 924. Although the degrader 920 can reduce theenergy, the accelerator 912 can be adjusted to output a lower beamenergy to reduce or eliminate the use of the degrader 920. The divisionof the target volumes 924, 930 into layers can be different or the same.And the target volume 930 can be treated similarly on a layer by layerbasis to the treatment of the target volume 924.

The treatment of the different target volumes 924, 930 on the samepatient may be substantially continuous, e.g., with the stop timebetween the two volumes being no longer than about 30 minutes or less,e.g., 25 minutes or less, 20 minutes or less, 15 minutes or less, 10minutes or less, 5 minutes or less, or 1 minute or less. As is explainedherein, the accelerator 912 can be mounted on a movable gantry and themovement of the gantry can move the accelerator to aim at differenttarget volumes. In some situations, the accelerator 912 can complete theenergy adjustment of the output beam 914 during the time the treatmentsystem makes adjustment (such as moving the gantry) after completing thetreatment of the target volume 924 and before starting treating thetarget volume 930. After the alignment of the accelerator and the targetvolume 930 is done, the treatment can begin with the adjusted, desiredbeam energy. Beam energy adjustment for different patients can also becompleted relatively efficiently. In some examples, all adjustments,including increasing/reducing beam energy and/or moving the gantry aredone within about 30 minutes, e.g., within about 25 minutes, withinabout 20 minutes, within about 15 minutes, within about 10 minutes orwithin about 5 minutes.

In the same layer of a target volume, an irradiation dose is applied bymoving the beam across the two-dimensional surface of the layer (whichis sometimes called scanning beam) using a scanning unit 916.Alternatively, the layer can be irradiated by passing the extracted beamthrough one or more scatterers of the scattering unit 16 (which issometimes called scattering beam).

Beam properties, such as energy and intensity, can be selected before atreatment or can be adjusted during the treatment by controlling theaccelerator 912 and/or other devices, such as the scanningunit/scatterer(s) 916, the degrader 920, and others not shown in thefigures. In this example implementation, as in the exampleimplementations described above, system 910 includes a controller 932,such as a computer, in communication with one or more devices in thesystem. Control can be based on results of the monitoring performed bythe one or more monitors 918, e.g., monitoring of the beam intensity,dose, beam location in the target volume, etc. Although the monitors 918are shown to be between the device 916 and the degrader 920, one or moremonitors can be placed at other appropriate locations along the beamirradiation path. Controller 932 can also store a treatment plan for oneor more target volumes (for the same patient and/or different patients).The treatment plan can be determined before the treatment starts and caninclude parameters, such as the shape of the target volume, the numberof irradiation layers, the irradiation dose for each layer, the numberof times each layer is irradiated, etc. The adjustment of a beamproperty within the system 910 can be performed based on the treatmentplan. Additional adjustment can be made during the treatment, e.g., whendeviation from the treatment plan is detected.

In some implementations, the accelerator 912 is configured to vary theenergy of the output particle beam by varying the magnetic field inwhich the particle beam is accelerated. In an example implementation,one or more sets of coils receives variable electrical current toproduce a variable magnetic field in the cavity. In some examples, oneset of coils receives a fixed electrical current, while one or moreother sets of coils receives a variable current so that the totalcurrent received by the coil sets varies. In some implementations, allsets of coils are superconducting. In other implementations, some setsof coils, such as the set for the fixed electrical current, aresuperconducting, while other sets of coils, such as the one or more setsfor the variable current, are non-superconducting. In some examples, allsets of coils are non-superconducting.

Generally, the magnitude of the magnetic field is scalable with themagnitude of the electrical current. Adjusting the total electriccurrent of the coils in a predetermined range can generate a magneticfield that varies in a corresponding, predetermined range. In someexamples, a continuous adjustment of the electrical current can lead toa continuous variation of the magnetic field and a continuous variationof the output beam energy. Alternatively, when the electrical currentapplied to the coils is adjusted in a non-continuous, step-wise manner,the magnetic field and the output beam energy also varies accordingly ina non-continuous (step-wise) manner. The scaling of the magnetic fieldto the current can allow the variation of the beam energy to be carriedout relatively precisely, although sometimes minor adjustment other thanthe input current may be performed.

In some implementations, to output particle beams having a variableenergy, the accelerator 912 is configured to apply RF voltages thatsweep over different ranges of frequencies, with each rangecorresponding to a different output beam energy. For example, if theaccelerator 912 is configured to produce three different output beamenergies, the RF voltage is capable of sweeping over three differentranges of frequencies. In another example, corresponding to continuousbeam energy variations, the RF voltage sweeps over frequency ranges thatcontinuously change. The different frequency ranges may have differentlower frequency and/or upper frequency boundaries.

The extraction channel may be configured to accommodate the range ofdifferent energies produced by the variable-energy particle accelerator.Particle beams having different energies can be extracted from theaccelerator 912 without altering the features of the regenerator that isused for extracting particle beams having a single energy. In otherimplementations, to accommodate the variable particle energy, theregenerator can be moved to disturb (e.g., change) different particleorbits in the manner described above and/or iron rods (magnetic shims)can be added or removed to change the magnetic field bump provided bythe regenerator. More specifically, different particle energies willtypically be at different particle orbits within the cavity. By movingthe regenerator in the manner described herein, it is possible tointercept a particle orbit at a specified energy and thereby provide thecorrect perturbation of that orbit so that particles at the specifiedenergy reach the extraction channel. In some implementations, movementof the regenerator (and/or addition/removal of magnetic shims) isperformed in real-time to match real-time changes in the particle beamenergy output by the accelerator. In other implementations, particleenergy is adjusted on a per-treatment basis, and movement of theregenerator (and/or addition/removal of magnetic shims) is performed inadvance of the treatment. In either case, movement of the regenerator(and/or addition/removal of magnetic shims) may be computer controlled.For example, a computer may control one or more motors that effectmovement of the regenerator and/or magnetic shims.

In some implementations, the regenerator is implemented using one ormore magnetic shims that are controllable to move to the appropriatelocation(s).

In some implementations, structure 715 (described above) is controlledto accommodate the different energies produced by the particleaccelerator. For example, structure 715 may be rotated so that anappropriate thickness intercepts a particle beam having a particularenergy. Structure 715 thus absorbs at least some of the energy in theparticle beam, enabling the particle beam to traverse the extractionchannel, as described above.

As an example, table 1 shows three example energy levels at whichexample accelerator 912 can output particle beams. The correspondingparameters for producing the three energy levels are also listed. Inthis regard, the magnet current refers to the total electrical currentapplied to the one or more coil sets in the accelerator 912; the maximumand minimum frequencies define the ranges in which the RF voltagesweeps; and “r” is the radial distance of a location to a center of thecavity in which the particles are accelerated.

TABLE 1 Examples of beam energies and respective parameters. MagneticBeam Magnet Maximum Minimum Field Magnetic Field Energy CurrentFrequency Frequency at r = 0 mm at r = 298 mm (MeV) (Amps) (MHz) (MHz)(Tesla) (Tesla) 250 1990 132 99 8.7 8.2 235 1920 128 97 8.4 8.0 211 1760120 93 7.9 7.5

Details that may be included in an example particle accelerator thatproduces charged particles having variable energies are described below.The accelerator can be a synchrocyclotron and the particles may beprotons. The particles output as pulsed beams. The energy of the beamoutput from the particle accelerator can be varied during the treatmentof one target volume in a patient, or between treatments of differenttarget volumes of the same patient or different patients. In someimplementations, settings of the accelerator are changed to vary thebeam energy when no beam (or particles) is output from the accelerator.The energy variation can be continuous or non-continuous over a desiredrange.

Referring to the example shown in FIG. 1, the particle accelerator(synchrocyclotron 502), which may be a variable-energy particleaccelerator like accelerator 912 described above, may be configured toparticle beams that have a variable energy. The range of the variableenergy can have an upper boundary that is about 200 MeV to about 300 MeVor higher, e.g., 200 MeV, about 205 MeV, about 210 MeV, about 215 MeV,about 220 MeV, about 225 MeV, about 230 MeV, about 235 MeV, about 240MeV, about 245 MeV, about 250 MeV, about 255 MeV, about 260 MeV, about265 MeV, about 270 MeV, about 275 MeV, about 280 MeV, about 285 MeV,about 290 MeV, about 295 MeV, or about 300 MeV or higher. The range canalso have a lower boundary that is about 100 MeV or lower to about 200MeV, e.g., about 100 MeV or lower, about 105 MeV, about 110 MeV, about115 MeV, about 120 MeV, about 125 MeV, about 130 MeV, about 135 MeV,about 140 MeV, about 145 MeV, about 150 MeV, about 155 MeV, about 160MeV, about 165 MeV, about 170 MeV, about 175 MeV, about 180 MeV, about185 MeV, about 190 MeV, about 195 MeV, about 200 MeV.

In some examples, the variation is non-continuous and the variation stepcan have a size of about 10 MeV or lower, about 15 MeV, about 20 MeV,about 25 MeV, about 30 MeV, about 35 MeV, about 40 MeV, about 45 MeV,about 50 MeV, about 55 MeV, about 60 MeV, about 65 MeV, about 70 MeV,about 75 MeV, or about 80 MeV or higher. Varying the energy by one stepsize can take no more than 30 minutes, e.g., about 25 minutes or less,about 20 minutes or less, about 15 minutes or less, about 10 minutes orless, about 5 minutes or less, about 1 minute or less, or about 30seconds or less. In other examples, the variation is continuous and theaccelerator can adjust the energy of the particle beam at a relativelyhigh rate, e.g., up to about 50 MeV per second, up to about 45 MeV persecond, up to about 40 MeV per second, up to about 35 MeV per second, upto about 30 MeV per second, up to about 25 MeV per second, up to about20 MeV per second, up to about 15 MeV per second, or up to about 10 MeVper second. The accelerator can be configured to adjust the particleenergy both continuously and non-continuously. For example, acombination of the continuous and non-continuous variation can be usedin a treatment of one target volume or in treatments of different targetvolumes. Flexible treatment planning and flexible treatment can beachieved.

A particle accelerator that outputs a particle beam having a variableenergy can provide accuracy in irradiation treatment and reduce thenumber of additional devices (other than the accelerator) used for thetreatment. For example, the use of degraders for changing the energy ofan output particle beam may be reduced or eliminated. The properties ofthe particle beam, such as intensity, focus, etc. can be controlled atthe particle accelerator and the particle beam can reach the targetvolume without substantial disturbance from the additional devices. Therelatively high variation rate of the beam energy can reduce treatmenttime and allow for efficient use of the treatment system.

In some implementations, the accelerator, such as the synchrocyclotron502 of FIG. 1, accelerates particles or particle beams to variableenergy levels by varying the magnetic field in the accelerator, whichcan be achieved by varying the electrical current applied to coils forgenerating the magnetic field. As shown in FIGS. 3, 4, 5, 6, and 7,example synchrocyclotron 10 (502 in FIG. 1) includes a magnet systemthat contains a particle source 90, a radiofrequency drive system 91,and a beam extraction system 38. FIG. 28 shows an example of a magnetsystem that may be used in a variable-energy accelerator. In thisexample implementation, the magnetic field established by the magnetsystem 1012 can vary by about 5% to about 35% of a maximum value of themagnetic field that two sets of coils 40 a and 40 b, and 42 a and 42 bare capable of generating. The magnetic field established by the magnetsystem has a shape appropriate to maintain focus of a contained protonbeam using a combination of the two sets of coils and a pair of shapedferromagnetic (e.g., low carbon steel) structures, examples of which areprovided above.

Each set of coils may be a split pair of annular coils to receiveelectrical current. In some situations, both sets of coils aresuperconducting. In other situations, only one set of the coils issuperconducting and the other set is non-superconducting or normalconducting (also discussed further below). It is also possible that bothsets of coils are non-superconducting. Suitable superconductingmaterials for use in the coils include niobium-3 tin (Nb3Sn) and/orniobium-titanium. Other normal conducting materials can include copper.Examples of the coil set constructions are described further below.

The two sets of coils can be electrically connected serially or inparallel. In some implementations, the total electrical current receivedby the two sets of coils can include about 2 million ampere turns toabout 10 million ampere turns, e.g., about 2.5 to about 7.5 millionampere turns or about 3.75 million ampere turns to about 5 millionampere turns. In some examples, one set of coils is configured toreceive a fixed (or constant) portion of the total variable electricalcurrent, while the other set of coils is configured to receive avariable portion of the total electrical current. The total electricalcurrent of the two coil sets varies with the variation of the current inone coil set. In other situations, the electrical current applied toboth sets of coils can vary. The variable total current in the two setsof coils can generate a magnetic field having a variable magnitude,which in turn varies the acceleration pathways of the particles andproduces particles having variable energies.

Generally, the magnitude of the magnetic field generated by the coil(s)is scalable to the magnitude of the total electrical current applied tothe coil(s). Based on the scalability, in some implementations, linearvariation of the magnetic field strength can be achieved by linearlychanging the total current of the coil sets. The total current can beadjusted at a relatively high rate that leads to a relatively high-rateadjustment of the magnetic field and the beam energy.

In the example reflected in Table 1 above, the ratio between values ofthe current and the magnetic field at the geometric center of the coilrings is: 1990:8.7 (approximately 228.7:1); 1920:8.4 (approximately228.6:1); 1760:7.9 (approximately 222.8:1). Accordingly, adjusting themagnitude of the total current applied to a superconducting coil(s) canproportionally (based on the ratio) adjust the magnitude of the magneticfield.

The scalability of the magnetic field to the total electrical current inthe example of Table 1 is also shown in the plot of FIG. 26, where BZ isthe magnetic field along the Z direction; and R is the radial distancemeasured from a geometric center of the coil rings along a directionperpendicular to the Z direction. The magnetic field has the highestvalue at the geometric center, and decreases as the distance Rincreases. The curves 1035, 1037 represent the magnetic field generatedby the same coil sets receiving different total electrical current: 1760Amperes and 1990 Amperes, respectively. The corresponding energies ofthe extracted particles are 211 MeV and 250 MeV, respectively. The twocurves 1035, 1037 have substantially the same shape and the differentparts of the curves 1035, 1037 are substantially parallel. As a result,either the curve 1035 or the curve 1037 can be linearly shifted tosubstantially match the other curve, indicating that the magnetic fieldis scalable to the total electrical current applied to the coil sets.

In some implementations, the scalability of the magnetic field to thetotal electrical current may not be perfect. For example, the ratiobetween the magnetic field and the current calculated based on theexample shown in table 1 is not constant. Also, as shown in FIG. 26, thelinear shift of one curve may not perfectly match the other curve. Insome implementations, the total current is applied to the coil setsunder the assumption of perfect scalability. The target magnetic field(under the assumption of perfect scalability) can be generated byadditionally altering the features, e.g., geometry, of the coils tocounteract the imperfection in the scalability. As one example,ferromagnetic (e.g., iron) rods (magnetic shims) can be inserted orremoved from one or both of the magnetic structures. The features of thecoils can be altered at a relatively high rate so that the rate of themagnetic field adjustment is not substantially affected as compared tothe situation in which the scalability is perfect and only theelectrical current needs to be adjusted. In the example of iron rods,the rods can be added or removed at the time scale of seconds orminutes, e.g., within 5 minutes, within 1 minute, less than 30 seconds,or less than 1 second.

In some implementations, settings of the accelerator, such as thecurrent applied to the coil sets, can be chosen based on the substantialscalability of the magnetic field to the total electrical current in thecoil sets.

Generally, to produce the total current that varies within a desiredrange, any combination of current applied to the two coil sets can beused. In an example, the coil set 42 a, 42 b can be configured toreceive a fixed electrical current corresponding to a lower boundary ofa desired range of the magnetic field. In the example shown in table 1,the fixed electrical current is 1760 Amperes. In addition, the coil set40 a, 40 b can be configured to receive a variable electrical currenthaving an upper boundary corresponding to a difference between an upperboundary and a lower boundary of the desired range of the magneticfield. In the example shown in table 1, the coil set 40 a, 40 b isconfigured to receive electrical current that varies between 0 Ampereand 230 Amperes.

In another example, the coil set 42 a, 42 b can be configured to receivea fixed electrical current corresponding to an upper boundary of adesired range of the magnetic field. In the example shown in table 1,the fixed current is 1990 Amperes. In addition, the coil set 40 a, 40 bcan be configured to receive a variable electrical current having anupper boundary corresponding to a difference between a lower boundaryand an upper boundary of the desired range of the magnetic field. In theexample shown in table 1, the coil set 40 a, 40 b is configured toreceive electrical current that varies between −230 Ampere and 0 Ampere.

The total variable magnetic field generated by the variable totalcurrent for accelerating the particles can have a maximum magnitudegreater than 4 Tesla, e.g., greater than 5 Tesla, greater than 6 Tesla,greater than 7 Tesla, greater than 8 Tesla, greater than 9 Tesla, orgreater than 10 Tesla, and up to about 20 Tesla or higher, e.g., up toabout 18 Tesla, up to about 15 Tesla, or up to about 12 Tesla. In someimplementations, variation of the total current in the coil sets canvary the magnetic field by about 0.2 Tesla to about 4.2 Tesla or more,e.g., about 0.2 Tesla to about 1.4 Tesla or about 0.6 Tesla to about 4.2Tesla. In some situations, the amount of variation of the magnetic fieldcan be proportional to the maximum magnitude.

FIG. 27 shows an example RF structure for sweeping the voltage on thedee plate 100 over an RF frequency range for each energy level of theparticle beam, and for varying the frequency range when the particlebeam energy is varied. The semicircular surfaces 103, 105 of the deeplate 100 are connected to an inner conductor 1300 and housed in anouter conductor 1302. The high voltage is applied to the dee plate 100from a power source (not shown, e.g., an oscillating voltage input)through a power coupling device 1304 that couples the power source tothe inner conductor. In some implementations, the coupling device 1304is positioned on the inner conductor 1300 to provide power transfer fromthe power source to the dee plate 100. In addition, the dee plate 100 iscoupled to variable reactive elements 1306, 1308 to perform the RFfrequency sweep for each particle energy level, and to change the RFfrequency range for different particle energy levels.

The variable reactive element 1306 can be a rotating capacitor that hasmultiple blades 1310 rotatable by a motor (not shown). By meshing orunmeshing the blades 1310 during each cycle of RF sweeping, thecapacitance of the RF structure changes, which in turn changes theresonant frequency of the RF structure. In some implementations, duringeach quarter cycle of the motor, the blades 1310 mesh with the eachother. The capacitance of the RF structure increases and the resonantfrequency decreases. The process reverses as the blades 1310 unmesh. Asa result, the power required to generate the high voltage applied to thedee plate 103 and necessary to accelerate the beam can be reduced by alarge factor. In some implementations, the shape of the blades 1310 ismachined to form the required dependence of resonant frequency on time.

The RF frequency generation is synchronized with the blade rotation bysensing the phase of the RF voltage in the resonator, keeping thealternating voltage on the dee plates close to the resonant frequency ofthe RF cavity. (The dummy dee is grounded and is not shown in FIG. 27).

The variable reactive element 1308 can be a capacitor formed by a plate1312 and a surface 1316 of the inner conductor 1300. The plate 1312 ismovable along a direction 1314 towards or away from the surface 1316.The capacitance of the capacitor changes as the distance D between theplate 1312 and the surface 1316 changes. For each frequency range to beswept for one particle energy, the distance D is at a set value, and tochange the frequency range, the plate 1312 is moved corresponding to thechange in the energy of the output beam.

In some implementations, the inner and outer conductors 1300, 1302 areformed of a metallic material, such as copper, aluminum, or silver. Theblades 1310 and the plate 1312 can also be formed of the same ordifferent metallic materials as the conductors 1300, 1302. The couplingdevice 1304 can be an electrical conductor. The variable reactiveelements 1306, 1308 can have other forms and can couple to the dee plate100 in other ways to perform the RF frequency sweep and the frequencyrange alteration. In some implementations, a single variable reactiveelement can be configured to perform the functions of both the variablereactive elements 1306, 1308. In other implementations, more than twovariable reactive elements can be used.

Any of the features described herein may be configured for use with avariable-energy particle accelerator, such as that described above.

Any two more of the foregoing implementations may be used in anappropriate combination to affect the energy of a particle beam in theextraction channel. Likewise, individual features of any two more of theforegoing implementations may be used in an appropriate combination forthe same purpose.

Elements of different implementations described herein may be combinedto form other implementations not specifically set forth above. Elementsmay be left out of the processes, systems, apparatus, etc., describedherein without adversely affecting their operation. Various separateelements may be combined into one or more individual elements to performthe functions described herein.

The example implementations described herein are not limited to use witha particle therapy system or to use with the example particle therapysystems described herein. Rather, the example implementations can beused in any appropriate system that directs accelerated particles to anoutput.

Additional information concerning the design of an exampleimplementation of a particle accelerator that may be used in a system asdescribed herein can be found in U.S. Provisional Application No.60/760,788, entitled “High-Field Superconducting Synchrocyclotron” andfiled Jan. 20, 2006; U.S. patent application Ser. No. 11/463,402,entitled “Magnet Structure For Particle Acceleration” and filed Aug. 9,2006; and U.S. Provisional Application No. 60/850,565, entitled“Cryogenic Vacuum Break Pneumatic Thermal Coupler” and filed Oct. 10,2006, all of which are incorporated herein by reference.

The following applications are incorporated by reference into thesubject application: the U.S. Provisional Application entitled“CONTROLLING INTENSITY OF A PARTICLE BEAM” (Application No. 61/707,466),the U.S. Provisional Application entitled “ADJUSTING ENERGY OF APARTICLE BEAM” (Application No. 61/707,515), the U.S. ProvisionalApplication entitled “ADJUSTING COIL POSITION” (Application No.61/707,548), the U.S. Provisional Application entitled “FOCUSING APARTICLE BEAM USING MAGNETIC FIELD FLUTTER” (Application No.61/707,572), the U.S. Provisional Application entitled “MAGNETIC FIELDREGENERATOR” (Application No. 61/707,590), the U.S. ProvisionalApplication entitled “FOCUSING A PARTICLE BEAM” (Application No.61/707,704), the U.S. Provisional Application entitled “CONTROLLINGPARTICLE THERAPY (Application No. 61/707,624), and the U.S. ProvisionalApplication entitled “CONTROL SYSTEM FOR A PARTICLE ACCELERATOR”(Application No. 61/707,645).

The following are also incorporated by reference into the subjectapplication: U.S. Pat. No. 7,728,311 which issued on Jun. 1, 2010, U.S.patent application Ser. No. 11/948,359 which was filed on Nov. 30, 2007,U.S. patent application Ser. No. 12/275,103 which was filed on Nov. 20,2008, U.S. patent application Ser. No. 11/948,662 which was filed onNov. 30, 2007, U.S. Provisional Application No. 60/991,454 which wasfiled on Nov. 30, 2007, U.S. Pat. No. 8,003,964 which issued on Aug. 23,2011, U.S. Pat. No. 7,208,748 which issued on Apr. 24, 2007, U.S. Pat.No. 7,402,963 which issued on Jul. 22, 2008, U.S. patent applicationSer. No. 13/148,000 filed Feb. 9, 2010, U.S. patent application Ser. No.11/937,573 filed on Nov. 9, 2007, U.S. patent application Ser. No.11/187,633, titled “A Programmable Radio Frequency Waveform Generatorfor a Synchrocyclotron,” filed Jul. 21, 2005, U.S. ProvisionalApplication No. 60/590,089, filed on Jul. 21, 2004, U.S. patentapplication Ser. No. 10/949,734, titled “A Programmable ParticleScatterer for Radiation Therapy Beam Formation”, filed Sep. 24, 2004,and U.S. Provisional Application No. 60/590,088, filed Jul. 21, 2005.

Any features of the subject application may be combined with one or moreappropriate features of the following: the U.S. Provisional Applicationentitled “CONTROLLING INTENSITY OF A PARTICLE BEAM” (Application No.61/707,466), the U.S. Provisional Application entitled “ADJUSTING ENERGYOF A PARTICLE BEAM” (Application No. 61/707,515), the U.S. ProvisionalApplication entitled “ADJUSTING COIL POSITION” (Application No.61/707,548), the U.S. Provisional Application entitled “FOCUSING APARTICLE BEAM USING MAGNETIC FIELD FLUTTER” (Application No.61/707,572), the U.S. Provisional Application entitled “MAGNETIC FIELDREGENERATOR” (Application No. 61/707,590), the U.S. ProvisionalApplication entitled “FOCUSING A PARTICLE BEAM” (Application No.61/707,704), the U.S. Provisional Application entitled “CONTROLLINGPARTICLE THERAPY (Application No. 61/707,624), and the U.S. ProvisionalApplication entitled “CONTROL SYSTEM FOR A PARTICLE ACCELERATOR”(Application No. 61/707,645), U.S. Pat. No. 7,728,311 which issued onJun. 1, 2010, U.S. patent application Ser. No. 11/948,359 which wasfiled on Nov. 30, 2007, U.S. patent application Ser. No. 12/275,103which was filed on Nov. 20, 2008, U.S. patent application Ser. No.11/948,662 which was filed on Nov. 30, 2007, U.S. ProvisionalApplication No. 60/991,454 which was filed on Nov. 30, 2007, U.S. patentapplication Ser. No. 13/907,601, which was filed on May 31, 2013, U.S.patent application Ser. No. 13/916,401, filed on Jun. 12, 2013, U.S.Pat. No. 8,003,964 which issued on Aug. 23, 2011, U.S. Pat. No.7,208,748 which issued on Apr. 24, 2007, U.S. Pat. No. 7,402,963 whichissued on Jul. 22, 2008, U.S. patent application Ser. No. 13/148,000filed Feb. 9, 2010, U.S. patent application Ser. No. 11/937,573 filed onNov. 9, 2007, U.S. patent application Ser. No. 11/187,633, titled “AProgrammable Radio Frequency Waveform Generator for a Synchrocyclotron,”filed Jul. 21, 2005, U.S. Provisional Application No. 60/590,089, filedon Jul. 21, 2004, U.S. patent application Ser. No. 10/949,734, titled “AProgrammable Particle Scatterer for Radiation Therapy Beam Formation”,filed Sep. 24, 2004, and U.S. Provisional Application No. 60/590,088,filed Jul. 21, 2005.

Except for the provisional application from which this patentapplication claims priority and the documents incorporated by referenceabove, no other documents are incorporated by reference into this patentapplication.

Other implementations not specifically described herein are also withinthe scope of the following claims.

What is claimed is:
 1. A particle accelerator comprising: a coil toprovide a magnetic field to a cavity; a particle source to provide aplasma column to the cavity; a voltage source to provide a radiofrequency (RF) voltage to the cavity to accelerate particles from theplasma column, the magnetic field causing particles accelerated from theplasma column to move orbitally within the cavity; an enclosurecontaining an extraction channel to receive the particles acceleratedfrom the plasma column and to output the received particles from thecavity; and a structure arranged proximate to the extraction channel tochange an energy level of the received particles.
 2. The particleaccelerator of claim 1, wherein the structure has multiple thicknesses;and wherein the structure is movable relative to the extraction channelto place one of the multiple thicknesses in a path of the receivedparticles.
 3. The particle accelerator of claim 2, wherein the structureis wheel-shaped and is rotatable within the extraction channel.
 4. Theparticle accelerator of claim 2, wherein the structure has variablethickness ranging from a maximum thickness to a minimum thickness. 5.The particle accelerator of claim 1, wherein the particle accelerator isrotatable relative to a fixed position; and wherein the particleaccelerator further comprises a control system to control movement ofthe structure based on a rotational position of the particleaccelerator.
 6. The particle accelerator of claim 1, further comprising:a regenerator to adjust the magnetic field within the cavity to therebychange successive orbits of the particles accelerated from the plasmacolumn so that, eventually, the particles output to the extractionchannel.
 7. The particle accelerator of claim 1, wherein the structurecomprises at least one of the following materials: beryllium, carbon andplastic.
 8. A proton therapy system comprising: the particle acceleratorof claim 1, wherein the particles comprise protons; and a gantry onwhich the particle accelerator is mounted, the gantry being rotatablerelative to a patient position; wherein protons are output essentiallydirectly from the particle accelerator to the patient position.
 9. Aparticle accelerator comprising: a coil to provide a magnetic field to acavity; a particle source to provide a plasma column to the cavity; avoltage source to provide a radio frequency (RF) voltage to the cavityto accelerate particles from the plasma column, the magnetic fieldcausing particles accelerated from the plasma column to move orbitallywithin the cavity; an enclosure containing an extraction channel toreceive the particles accelerated from the plasma column and to outputthe received particles from the cavity; and a regenerator to adjust themagnetic field within the cavity to thereby change successive orbits ofthe particles accelerated from the plasma column so that, eventually,the particles output to the extraction channel, the regenerator beingmovable within the cavity relative to orbits of the particles.
 10. Theparticle accelerator of claim 9, wherein the regenerator is configuredto move radially relative to an approximate center of the cavity. 11.The particle accelerator of claim 10, further comprising: an actuator tomove the regenerator in response to a control signal.
 12. The particleaccelerator of claim 11, wherein the particle accelerator is rotatablerelative to a fixed position; and wherein the particle acceleratorfurther comprises a control system to generate the control signal tocontrol movement of the regenerator based on a rotational position ofthe particle accelerator.
 13. The particle accelerator of claim 9,wherein the regenerator comprises a ferromagnetic material.
 14. A protontherapy system comprising: the particle accelerator of claim 9, whereinthe particles comprise protons; and a gantry on which the particleaccelerator is mounted, the gantry being rotatable relative to a patientposition; wherein protons are output essentially directly from theparticle accelerator to the patient position.
 15. A particle acceleratorcomprising: a coil to provide a magnetic field to a cavity; a particlesource to provide a plasma column to the cavity; a voltage source toprovide a radio frequency (RF) voltage to the cavity to accelerateparticles from the plasma column, the magnetic field causing particlesaccelerated from the plasma column to move orbitally within the cavity;an enclosure containing an extraction channel to receive the particlesaccelerated from the plasma column and to output the received particlesfrom the cavity; and a regenerator to adjust the magnetic field withinthe cavity to thereby change successive orbits of the particlesaccelerated from the plasma column so that, eventually, the particlesoutput to the extraction channel; wherein the enclosure comprisemagnetic structures, at least one of the magnetic structures having aslot therein, the slot containing a magnetic shim that is ferromagneticand movable within the slot, the magnetic shim being movable relative tothe regenerator to affect an amount by which the regenerator adjusts themagnetic field.
 16. The particle accelerator of claim 15, wherein the atleast one of the magnetic structures has multiple slots therein, eachslot containing a magnetic shim that is ferromagnetic and movable withinthe slot, each magnetic shim being movable relative to the regeneratorto affect an amount by which the regenerator adjusts the magnetic field.17. The particle accelerator of claim 15, wherein the particleaccelerator is rotatable relative to a fixed position; and wherein theparticle accelerator further comprises a control system to generate acontrol signal to control movement of the magnetic shim based on arotational position of the particle accelerator.
 18. The particleaccelerator of claim 15, wherein the magnetic shim comprises anelectromagnet.
 19. A proton therapy system comprising: the particleaccelerator of claim 15, wherein the particles comprise protons; and agantry on which the particle accelerator is mounted, the gantry beingrotatable relative to a patient position; wherein protons are outputessentially directly from the particle accelerator to the patientposition.
 20. A particle accelerator comprising: a cryostat comprising asuperconducting coil, the superconducting coil conducting a current thatgenerates a magnetic field; magnetic structures adjacent to thecryostat, the cryostat being attached to the magnetic structures, themagnetic structures containing a cavity; a particle source to provide aplasma column to the cavity; a voltage source to provide a radiofrequency (RF) voltage to the cavity to accelerate particles from theplasma column, the magnetic field causing particles accelerated from theplasma column to move orbitally within the cavity; an extraction channelto receive the particles accelerated from the plasma column and tooutput the received particles from the cavity; and an actuator that iscontrollable to move the cryostat relative to the magnetic structures.21. The particle accelerator of claim 20, wherein the particleaccelerator is rotatable relative to a fixed position; and wherein theparticle accelerator further comprises a control system to generate acontrol signal to control the actuator based on a rotational position ofthe particle accelerator.
 22. The particle accelerator of claim 21,wherein the actuator is controlled to control movement of the cryostatso as to compensate for effects of gravity on the superconducting coil.23. A proton therapy system comprising: the particle accelerator ofclaim 20, wherein the particles comprise protons; and a gantry on whichthe particle accelerator is mounted, the gantry being rotatable relativeto a patient position; wherein protons are output essentially directlyfrom the particle accelerator to the patient position.
 24. Avariable-energy particle accelerator comprising: magnetic structuresdefining a cavity in which particles are accelerated for output as aparticle beam that has a selected energy from among a range of energies;an extraction channel to receive the particle beam; and a structureproximate to the extraction channel to intercept the particle beam priorto the particle beam entering the extraction channel, the structurebeing movable based on the selected energy, the structure to absorb atleast some energy of the particle beam prior to the particle beamentering the extraction channel.
 25. The variable-energy particleaccelerator of claim 24, wherein the structure comprises a wheel havingvarying thickness, where different thicknesses are capable of absorbingdifferent amounts of energy.
 26. The variable-energy particleaccelerator of claim 24, further comprising: a magnetic regenerator toimplement a magnetic field bump at a particle orbit that corresponds tothe selected energy; wherein the magnetic regenerator is movable basedon movement of the variable-energy particle accelerator.
 27. Thevariable-energy particle accelerator of claim 24, further comprising: amagnetic regenerator to implement a magnetic field bump at a particleorbit that corresponds to the selected energy; wherein the magneticregenerator is movable to intercept a particle orbit having the selectedenergy.